The present invention pertains to the art of electrical sensing devices and, more particularly, to a method and apparatus for sensing a time varying electrical current passing through an ion channel located in a membrane. A membrane typically surrounds biologic cells, with the membrane having ion channels located therein. These ion channels are made of protein structures that regulate the passage of various ions into and out of the cell. These ion channels are important to study due to their wide variety of biological functions. In the past, such ion channels were studied by clamping a portion or patch of a cell membrane and measuring changes in its electrical properties when the membrane and ion channels were exposed to various analytes which altered the ion channel's behavior. Unfortunately, there are great difficulties in positioning a cell for such measurement. In order to make ion channel studies more efficient, people have modeled the cell membrane by forming a lipid bilayer membrane and inserting therein proteins that act as ion channels. Such an artificial membrane may then be used in experiments rather than an actual cell membrane.
In a manner known in the art, measurements of ionic current passing through an ion channel or protein pore, nanochannel or other aperture of a membrane maintained in an electrolyte have been made using first and second electrodes in resistive electrical contact with the electrolyte. The first electrode is maintained in a first or bath volume of electrolyte, while the second electrode is maintained in a second or sensing volume of electrolyte. A voltage differential is maintained between the first and second electrodes establishing an electrical field. At an interface between the first (resistive) electrode and an ionic solution, an oxidation-reduction (redox) reaction must occur for a charge to transfer between the first electrode and the solution. In the oxidation part of the reaction, atoms of the first electrode enter the solution as cations. The cations move through the solution under the influence of diffusion and the electrical field and are deposited at the second electrode as the cations are reduced.
The interaction between the resistive electrode and the solution produces a concentration gradient of the ions in solution near the surface of the first electrode. The gradient causes a half-cell potential to be established between the first electrode and the bulk of the electrolyte. If a current flows, these potentials can be altered causing an over-voltage to appear. The over-voltage results from an alteration in a charge distribution of the solution in contact with the first and second electrodes, creating a polarization effect. The resulting charge distribution can cause a significant measurement artifact if either the first or second electrodes move relative to the solution for any reason.
In order to minimize these well-known problems with resistive electrodes, prior measurement apparatuses have positioned the measurement volume of interest away from a region of variable concentration around the electrodes. In suspended membrane geometry, volumes having dimensions in the order of 1 cm×1 cm×1 cm are typically utilized on either side of the membrane. A wire electrode is immersed in each volume at a distance in the order of millimeters from an active area of the channel (or pore etc.). At this distance scale concentration gradient effects are negligible. However, it should be readily apparent that as the scale of the measurement apparatus is reduced, it becomes increasingly difficult to spatially separate the area with concentration gradients from the active region of interest. Indeed, in the limiting case of a supported membrane, a geometry involving a membrane deposited directly on an electrode, or separated from it by a thin (1 nm) layer, it has not yet proven possible to record a signal from a single channel.
In addition, as the overall scale of the apparatus is reduced, the volume is also reduced and the duration over which the second electrode can maintain an ionic current before the electrode is fully dissolved is reduced accordingly. Traditional patch clamp type experiments are limited to approximately one hour due to limitations in the lifetime of various aspects of the system. However, electrode degradation is not usually a limiting factor. In any case, for new applications that seek to study long-term effects and for systems with micrometer or nanometer scale electrodes, the lifetime of a resistively coupled electrode could be a limiting factor. A similar problem occurs due to the build up of ions from solution. If the region around the electrode is limited to nanometers by, for example, the presence of a supported membrane, then the deposition of even a nanometer of atoms from solution can present a significant problem.
A capacitive electrode does not suffer redox and concentration related problems as the electrode is insulated from the solution. Therefore, no ionic reaction occurs at the electrode. However, the capacitive electrode does produce a potential in the electrolyte. A capacitive electrode couples to the electrolyte by virtue of its mutual capacitance to the electrolyte. This potential induces ions to flow in the body of the electrolyte just as if a resistive electrode coupled the potential. An oscillating ionic current is maintained in the electrolyte by a displacement current induced in insulation around the capacitive electrode. In an identical manner, a capacitive electrode can also be used to measure the potential of an ionic fluid.
Despite these benefits, capacitive electrodes have not previously been used to measure potentials or currents in electrolytes. The reason is that existing biopotential electrodes have been adequate for experimental scale geometries utilized to date and have the benefit of being DC coupled. In addition, although the benefit of capacitive electrodes increases with reduced apparatus size, capacitive electrodes themselves become more difficult to use. That is, as the electrodes are made smaller, the capacitance of the electrode is reduced to a very small level. For example, for a 10 μm×10 μm electrode that might be used in a chip scale sensor, the capacitance of the electrode is in the order of less than 1 pF. At low frequency, a capacitance in the order of 1 pF represents very high impedance. Coupling an amplifier efficiently to such a high impedance source while maintaining low input noise levels and removing low frequency drift is traditionally a difficult problem.
New ways to couple to a very small, purely capacitive source have been taught by U.S. Pat. No. 6,686,800 B2. New systems that utilize such capacitive sensing to measure electrophysiological signals such as the human electrocardiogram (ECG) and electroencephalogram (EEG) are known in the art. In these cases, the capacitance of the region used to sense the potential was in the range of 10 pF to less than 1 pF. Other prior art arrangements used a capacitive method to measure the potential of a cell. The method employed an in situ transistor with the cell being deposited directly on a gate of a transistor. The internal potential of the cell and the potential of a cleft region, i.e., a small region of fluid between the cell and the upper surface of the transistor, were coupled into the transistor to produce a measurable signal.
In prior measurement arrangements that employ low-capacitance sensors, the variable of interest was the electric potential produced within a heart, brain, or other cell. In the case of the cell, the potential of the cleft region between the cell outer surface and the point of measurement was, in most cases, found to obscure and dominate the cell potential. The potential of the cleft region is determined by a combination of capacitive coupling to the internal potential of the cell, ionic current that flows through channels in a portion of the cell that faces the cleft region, and a resistive coupling through the electrolyte in the cleft region to the bath that maintains the cell. Owing to variations in spacing the cell from the electrode (i.e. the height of the cleft region) and in the local properties of the cell membrane, there is considerable, uncontrollable variation in the coupling of the cell potential to the transistor.
As a way to better control the cleft region, the prior art teaches attaching a lipid vesicle to a transistor instead of attaching the cell to the transistor. Because the lipid vesicle lacks an outer coating of proteins and oligo-saccharides, found around most cells, the lipid vesicle forms a cleft region having a lower, and a more reproducible height. However, as for the cell, there is a continuous fluid path from the electrolyte in the cleft region to the reservoir that contains the bulk of the electrolyte. Accordingly, there is always a conducting path from the cleft region to the electrode in the reservoir. The resistance of the conducting path is not a property of the membrane, but a difficult-to-control variable that depends on proper adhesion of the lipid vesicle to the transistor.
Other capacitive sensing configurations teach depositing a black lipid membrane containing gramicidin channels over an insulating groove provided in a silicon substrate. A linear array of transistors is fabricated at a bottom surface of the insulating groove. Electrodes are fabricated within the insulating groove at each end of the linear array in order to drive a current along a length of the array. A change in density of open channels was determined by a change in a voltage profile along the length of the groove.
Based on the above, there still exists a need for sensing time varying current passing through a membrane. More specifically, there exists a need for an apparatus that establishes general measurement geometry and associated electronic biasing techniques to enable a capacitive sensor to measure an ionic current passed by an ion channel or protein pore of a membrane.